The present embodiments relate to a counting digital x-ray detector and to a method for recording x-ray images.
X-ray systems are used for imaging for diagnostic examination and for interventional procedures in cardiology, radiology and surgery, for example. X-ray systems 16, as shown in FIG. 1, have an x-ray tube 18 and an x-ray detector 17, disposed together on a C-arm 19, for example, a high-voltage generator for generation of the tube voltage, an imaging system 21 (e.g., inclusive of at least one monitor 22), a system control unit 20, and a patient table 23. Systems with two planes (e.g., two C-arms) are also used in interventional radiology. Flat-panel x-ray detectors may be used as x-ray detectors in many areas of medical x-ray diagnostics and intervention (e.g., in radiography, interventional radiology, cardio angiography, but also in therapy for imaging within the context of checking and radiation therapy planning or mammography).
Current flat-panel detectors may be integrating detectors and are based primarily on scintillators, the light of which is converted in matrices of photodiodes into electrical charge. These may be be read out row-by-row via active control elements. FIG. 2 shows the basic structure of an indirect converting flat-panel x-ray detector currently used, having a scintillator 10, an active read-out matrix 11 made of amorphous silicon with a plurality of pixel elements 12 (e.g., with photodiode 13 and switching element 14) and activation and read-out electronics 15 (see, e.g., M. Spahn, “Flat detectors and their clinical applications,” Eur Radiol, (2005), 15: 1934-1947). Depending on the beam quality, the quanta efficiency for a scintillator made of CsJ with a layer thickness of, for example, 600 μm, depending on beam quality, lies between around 50% and 80% (see, e.g., M. Spahn, “Flat detectors and their clinical applications,” Eur Radiol (2005), 15: 1934-1947). The local-frequency-dependent detective quantum efficiency DQE(f) is delimited upwards by this and for typical pixel sizes of, for example, 150 μm to 200 μm and, for the local frequencies of 1 to 2 lp/mm of interest for the application, lies well below this figure. To make new applications (e.g., dual-energy, material-separation) possible, but also to increase the quanta efficiency further, the potential of counting detectors or energy-discriminating counting detectors mainly based on direct-converting materials (e.g., CdTe or CdZTe=CZT) and contacted application specific integrated circuits (ASICs) (e.g., implemented in CMOS technology) is increasingly being investigated.
A possible layout of such counting detectors is presented in FIG. 3. X-radiation is converted in direct converter 24 (e.g., CdTe or CZT), and the charge carrier pairs generated are separated via an electrical field that is created by a common top electrode 26 and a pixel electrode 25. The charge creates a charging pulse in one of the pixel electrodes 26 configured in a pixel shape of the ASIC 27, the height of which corresponds to the energy of the x-ray quantum and which, if lying above a defined threshold value, is registered as a count event. The threshold value serves to distinguish between an actual event and electronic noise or also to suppress k fluorescence photons, for example, in order to avoid multiple counting. The ASIC 27, a corresponding section of the direct converter 24 and a coupling between direct converter 24 and ASIC 27 (e.g., connected to each other by bump bonds 36 in direct-converting detectors) each form the detector module 35 with a plurality of pixel elements 12. The ASIC 27 is disposed on a substrate 37 and is connected to peripheral electronics 38. A detector module may also have one or more ASICs and one or more part sections of a direct converter, selected as required.
FIG. 4 shows the general schematic of a counting pixel element 12. The electrical charge is collected at a pixel electrode 28 via the electrical field applied and amplified with the aid of a charge amplifier 29 and a feedback capacitor 40. In addition, the pulse shape may be adapted in a shaper (e.g., filter) at the output (not shown). An event is then counted by a digital memory unit (counter) 33 being incremented by one if the output signal lies above a threshold value that may be set. This is set via a discriminator 31. The threshold value may also be specified as a fixed analog value, but is generally applied via a digital-to-analog converter DAC 32, for example, and is thus able to be set variably in a certain range. The threshold value may either be set for each pixel locally (e.g., by the discriminator 31 and the DAC 32, as shown) or is also able to be set globally in the x-ray detector for all pixel elements. Subsequently, there may be a read out via an activation and read-out unit 38. FIG. 6 shows a corresponding schematic for an overall array of counting pixel elements 12 (e.g., 100×100 pixel elements each of 180 μm). In this example, the array would have a size of 1.8×1.8 cm2. For large-surface detectors (e.g., 20×30 cm2), a number of detector modules 35 are combined (e.g., 11×17 would roughly produce this surface) and connected via the common peripheral electronics. Through silicon via (TSV) technology is used, for example, for the connection between ASIC and peripheral electronics in order to provide a four-sided arrangement of the modules as close to one another as possible.
In the case of counting and energy-discriminating x-ray detectors, two, three or more threshold values are introduced, and the level of the charge pulse, corresponding to the predefined threshold values (e.g., discriminator threshold values) is classified into one or more of the digital memory units (e.g., counters). The x-ray quanta counted in a particular energy range may be obtained by forming the difference of the counter contents of two corresponding counters. The discriminators are able to be set, for example, with the aid of digital-to-analog converters (DACs) for the detector module as a whole or for pixels within given limits or ranges. The counter contents of the pixel elements are read out module-by-module via a corresponding read-out unit. The read-out process uses a certain time during which further counting may not be performed without errors.
For x-ray quantum energies above the k edge of the respective detector material used (e.g., 27 keV for Cd, 32 keV for Te), k fluorescence dominates in the photo effect. As well as the photoelectron, a k fluorescence photon 42 may be re-emitted during the absorption of an x-ray quantum 41 that has somewhat less than the energy of the k edge (e.g., difference of the binding energies of the k shell and the shell from which the promoted electron originates), shown in FIG. 5. On account of the non-negligible average free wavelength of the k fluorescence photon 42 (e.g., approximately 120 μm for Cd and 66 μm for Te), three cases may now occur: (i) reabsorbed into the same pixel element as the primary photon; (ii) reabsorbed into a neighboring pixel element; or (iii) leaves the detector material completely.
The case of re-absorption into a neighboring pixel element (e.g., more likely for the x-ray energies typical in medical imaging than the case of not being absorbed in the detector material at all) now results in two count events occurring and the energy of the primary quantum being distributed over both pixel elements. Thus, both the counting rate and also the energies detected in each case are incorrect. The smaller the layout of the pixel elements, the more likely this case is to occur. For pixel sizes as are currently used in angiography, this is already a significant problem.
Other effects such as what is known as charge sharing (e.g., the distribution of the charge cloud that is generated at the edge of a pixel element) on this and at least one further neighboring pixel element may lead to similar effects such as k escape (e.g., k-fluorescence); depending on threshold value settings, pixel size, high voltage or field distribution in the detector material, absorption location, to multiple counting and a division of the energy between two or more pixel elements.
For counting and energy resolving x-ray detectors, it is important for the correct number of an x-ray quanta and the correct energy to be registered. To bring this about, there is the approach of resolving this problem by checking the coincidence (e.g., the essentially simultaneous occurrence) of two events in neighboring pixel elements. If two or more events thus essentially occur simultaneously in neighboring pixel elements, it may be assumed that the same event is involved. The probability of true coincidence (i.e., the same event) or false coincidence (i.e., the random simultaneous arrival of two primary quanta in neighboring pixel elements) being involved is dependent on a number of factors including x-ray flux, a temporal pulse width of the charge pulse or the size of the pixel elements. The greater the x-ray flux, the greater the probability of different events occurring in neighboring pixel elements (e.g., false coincidences).
In order to resolve the problem described above, next neighbor coincidence circuits may be used. In such cases, a check is made as to whether, starting from a pixel element, the direct neighboring pixel elements have likewise detected a count event above a given threshold (e.g., discriminator threshold). If this is the case, the events are only counted once, and the other signals are discarded. If the detector design is additionally energy-discriminating, then the total signal from the central and neighboring pixel elements is additionally combined, and this is sorted into the corresponding pixel element (e.g., the element with the highest signal) or into the counter or counters of the corresponding pixel element of which the threshold values are below the combined signal. Thus, for example, the pixel design has four counters with threshold values that correspond to energies of 25 keV, 45 keV, 65 keV and 85 keV. If the signal detected and combined according to coincidence is, for example, 50 keV, the signal will be classified (counted) into the counters with the threshold values at 25 keV and 45 keV. In general, the signal combination will be performed in the analog range in order to reconstruct the energy in the best possible manner. A digital coincidence circuit may however also be provided. Since a counter event always arises above the defined threshold values, energy intervals are generated by differentiation (e.g., in the above example, the energy interval from 45 to 65 keV by subtraction of the count events in the counter with threshold value 65 keV and that with the threshold value 45 keV). What has been described above is applicable for simple discriminators that only define one lower threshold value as a threshold. In principle, window discriminators that have both a lower and also an upper threshold value are also able to be used. Differentiation, as described above, is not then necessary.
Known coincidence circuits are shown in FIGS. 7 and 8, where circuit technology details are not shown. FIG. 7 shows a coincidence circuit in which the analog signals of the central pixel element 12.1 and the analog coincidence signals 45 of all direct neighboring pixel elements 12.2 are summed as analog values by a summation unit 44 and are then discriminated. If the sum signal lies above the discriminator threshold value 32 of the discriminator 31, then an event is counted. At the same time (not shown by the circuit diagram), no events are counted in the neighboring pixel elements 12.2. The figure also does not show that a number of discriminators with different threshold values and a number of counters may be present to discriminate different energies and count corresponding events. In addition, an additional discriminator that is to be exceeded to supply the analog signal to a summation may also be present in each pixel element. FIG. 8 shows a variant in which the coincidence circuit is realized as a digital circuit.
Suitable direct converters, which make high signals and count rates possible, such as, for example, CdTe or CZT may only be manufactured with known methods in small surfaces (e.g., of 2×2 cm2 or 3×3 cm2). ASICs with a complex pixel structure, as is needed for counting detectors, are currently only able to be manufactured with viable yields in small surfaces (e.g., similar dimensions as the detector material such as 2×2 cm2 or 3×3 cm2, possibly up to 2×8 cm2 or 3×6 cm2, so that in these examples four 2×2 cm2 or two 3×3 cm2 direct converter semiconductor pieces may be accommodated on the corresponding ASICs). Such detector modules are small compared to the overall size of an x-ray detector needed for applications in angiography, for example (e.g., 20×20 cm2 or 30×40 cm2).
The relatively small module surfaces provide that there are many pixel elements that occur at the edge or at a corner of the detector module. Their behavior differs from the behavior of pixel elements lying in the center, for example, because of the following effects: (i) k-fluorescence photons have a higher probability with edge pixels of escaping from the detector material than pixels lying in the center; (ii) the active surface of edge and corner pixel elements is frequently smaller than that of central pixel elements (e.g., on account of a guard ring at the edge of the detector material or in order to arrange the detector modules next to one another without loss of one or more pixel element rows or columns; and (iii) the field distribution at edge or corner pixel and thus possibly the charge collection efficiency may differ from the central pixel elements. Because of this and further effects, the response behavior of edge or corner pixel elements may differ from the response behavior of central pixel elements. In FIGS. 9 to 11, simplified pixel elements 12 with a coincidence circuit with a summation unit 44 are shown having different numbers of neighboring pixel elements. The rest of the circuit elements have been omitted for the sake of simplicity. The pixel element of FIG. 9 has eight neighboring pixel elements, which have been included in the coincidence circuit, the pixel element of FIG. 10 has five neighboring pixel elements, and the pixel element of FIG. 11 has three neighboring pixel elements.
Known coincidence circuits (e.g., next-neighbor coincidence) have at least two types of problems in conjunction with counting, direct-converting detector modules. At high counting rates, it is no longer possible to discriminate between true and false coincidences. Also, for pixel elements at the edges of the detector modules, connection to the pixel elements at the edges of the neighboring detector module is not able to be realized in practice, so that next-neighbor information is only partly available.